It is recognized that progression of such diseases as cancer and atherosclerosis, for example, and other debilitating disorders including neurodegenerative disease and osteoarthritis, is accompanied by changes in stiffness of biological tissue. Recent advances in the field of mechanobiology establish that these changes in the stiffness of the extra-cellular matrix (ECM) are not merely passive consequences of earlier causal events, but may in turn influence the behavior of tissue cells, thereby possibly further exacerbating the disease. The biological cells are mechanosensitive in that they feel, perceive, and respond to the mechanical properties of their ECM microenvironment. For example, a cell senses stiffness by exerting tension as it anchors and pulls on the ECM via focal adhesion sites that involve transmembrane integrins and a network of intracellular mechanosensory proteins. Mechanical cues received from the ECM are relayed and translated by intracellular signaling pathways that, in turn, influence cell morphology, differentiation, proliferation, contractility and elasticity. Behavior of the cells that have been altered affects a dynamic balance between the ECM production and break down, thereby causing the ECM stiffness to be changed further. As a result, a positive feedback loop is established with consequences that are sometimes detrimental to the cell's health. For example, the altered ECM stiffness can induce epithelial tumor progression, switch on the malignant phenotype in tumor cells, cause smooth muscle cell proliferation in atherosclerosis, enhance the angiogenesis potential of endothelial cells, initiate calcium deposition by interstitial cells in cardiac valves, modulate stem cell differentiation, and induce cell apoptosis. The cellular response was shown to be regulable via tuning the ECM mechanical properties to values comparable with those of a normal tissue.
Changes in the mechanical properties of the ECM may provide the early detectable signs of the disease onset that likely precede aberrant intracellular signals. Moreover, by engineering the ECM mechanical properties it may be possible to reverse the progress of the disease. Therefore, the capability to measure and monitor minute changes in the ECM stiffness at the size scale sensed by cells (referred to herein as cellular spatial scale) is vital in advancing current understanding in mechanobiology and may, quite possibly, enable not only the detection of the initial onset of a number of critical diseases but also the guidance of an early therapeutic intervention in case of such diseases.
The currently used systems and method are adapted to in vitro studies that evaluate the impact of global (or bulk) ECM mechanical properties on condition of the cells. In contradistinction, however, the biological cells probe the stiffness of their local microenvironment on a substantially smaller scale, via micron-sized focal adhesions and, due to tissue heterogeneities and matrix remodeling, the ECM micromechanical environment that a cell perceives is vastly different from the bulk mechanical environment. The majority of the hypotheses in mechanobiology, generated from experiments in monolayer cell models, fail to recapitulate the complex three-dimensional (3D) environment that a cell experiences in vivo. It is well established that cellular behavior is profoundly different in 3D models where the influence of the ECM composition and stiffness is far more complicated compared to the two-dimensional (2D) monolayer models. Accordingly, there remains a question of how mechanobiological relationships translate into biologically relevant 3D disease models and in clinically relevant systems in vivo. However, no means exists today that enable measurements of the ECM stiffness in 3D at microscopic size scales relevant to the microenvironment of a biological cell.
It is increasingly appreciated that the ECM, also known as the tissue scaffolding, not only provides mechanical stability and tissue organization, but also imparts critical biochemical and biomechanical cues. Such cues actively direct cell growth, survival, and migration, and govern vascularization and immune responsiveness during embryonic development, homeostasis maintenance, and disease progression. ECM biomechanical properties play a prominent role in neoplastic transformation and metastatic progression, which accounts for more than 90% of cancer-associated mortality and morbidity. In addition, these micromechanical cues are directly implicated in many fibrotic diseases, such as idiopathic pulmonary fibrosis (IPF), systemic sclerosis (SS), liver cirrhosis, and atherosclerosis, which are responsible for over 45% of deaths in the industrialized countries.
For instance, in breast cancer, the biophysical and biochemical cues from the tumor-associated ECM reinforce and fuel the progression of the neoplasia by promoting of the hallmarks of cancer. The increased ECM stiffness, sensed by focal adhesions, can activate integrins, promote focal adhesion assembly, and/or stimulate the mechano-sensory pathways. This activation increases cytoskeletal tension via acto-myosin contractility. This, in turn, can accelerate the secretion, deposition, and cross linking of ECM macromolecules by host stromal cells, and results in further increases in ECM stiffness, completing a self-enforcing vicious cycle. In addition, interference between cytoskeletal tension and the epidermal growth-factor receptor (EGFR) pathways can result in increased proliferation of cancer cells. The biophysical properties of ECM are also known to promote pro-migratory trails by inducing epithelial-mesenchymal transition. By increasing the VEGF signaling, ECM stiffening can also provoke angiogenesis. Further, increased ECM density and rigidity can amplify the interstitial fluid pressure and hamper drug delivery, promoting biophysical drug resistance.
Apart from solid tumors, the micromechanical signature of ECM is also directly implicated in the development of numerous fibrotic disorders. One such disease is IPF, a devastatingly progressive fibrosis of the lungs that destroys the normal alveolar structures and impairs oxygen transfer to the blood stream. The etiology of IPF is not well understood, yet ECM micromechanics are believed to be directly involved in the pathogenesis of IPF. Stiffening of ECM activates the mechano-sensory circuitry of fibroblasts, in turn promoting contractility, proliferation, acquired resistance to apoptosis, and differentiation to contractile myofibroblast phenotype. The subsequent increased collagen synthesis and accumulation continuously translates the mechanical stimuli into fibrogenic signals and vice-versa. Despite the ubiquitous role of ECM micro-mechanics in normal lung development and pulmonary disorders such as IPF, our current understanding of ECM mechanics is limited. The currently available models of lung biomechanics are overly simplified, ignoring the local tissue heterogeneity at cellular scales.
Irregular ECM micromechanics are also involved in the initiation and progression of systemic sclerosis (SS) or scleroderma. The limited cutaneous scleroderma affects only the skin on the face and limbs, whereas the diffuse type is likely to damage internal organs such as kidneys, heart, lungs, and digestive tract. Currently, the cause of SS is unknown and no pharmacotherapy is available for this deadly disease. However, it is believed that onset of SS is triggered by an injury followed by an aberrant wound healing response which involves activation of inflammatory pathways, accumulation of collagen and other fibrous proteins, and ECM stiffening within the dermis. As in IPF, the host stromal cells within the ECM transform into persistent myofibroblasts. The subsequent excessive deposition of ECM creates a feed-forward loop with catastrophic consequences. The ability to quantify and map ECM micro-mechanical properties is crucial for finding novel therapies that target the ECM.
Evaluating the ECM viscoelastic properties at micro-scale, as perceived by the cells, may help us understand the mechanical regulation of many diseases. In addition, it may provide a strong diagnostic tool for staging the disease, tailoring effective and personalized therapeutic strategies, and monitoring the efficacy of the treatments. It may also open new prognostic and therapeutic avenues that target ECM mechanical properties to reverse the course of the disease and regress its progression.
Apart from differentiating between natural tissue scaffolds in healthy and disease states, diagnosis of malignancies, understanding the etiology and pathogeneses of multiple conditions, and devising therapeutic approaches based on regulating the mechanical properties of tissues, the ability to evaluate the viscoelastic properties at multiple spatial scales is invaluable for design and development of synthetic tissue scaffolds, biomaterials, and hydrogels. This is because biomaterials and hydrogels are increasingly used in tissue engineering, regenerative medicine, drug-delivery, and mechanobiology research owing to their unique properties, including biocompatibility, tunable compliance, deformability and stress resilience. To fully integrate into biological systems, biomaterials and tissue scaffoldings often exhibit distinct viscoelastic properties at different length scales. The macro-scale viscoelastic endurance of biomimetic scaffolds enables them to withstand the physiological and hemodynamic loads. The micro-scale compliance, on the other hand, confers mechanical stimuli that direct cellular growth and differentiations and control molecular dynamics such as the gas and nutrients exchange, and drug-release. The growing demand for design, development, quality control, and performance monitoring of biomimetic constructs calls for novel tools capable of evaluating the viscoelasticity of biomaterials, in their native state, at different deformation rates and multiple length-scales, without manipulation.
The bulk mechanical properties of materials are often quantified by the frequency dependent shear viscoelastic modulus, G*(ω)=G′(ω)+i G″(ω). Here, G* is the shear viscoelastic modulus and ω is the oscillation frequency of the loading condition. Moreover, G′, the real part of G*, is the elastic modulus, representing the solid-like behavior of the sample, and G″(ω), the imaginary part of G*, is the viscous modulus, characterizing the fluid-like trait of the specimen. Traditionally, G*(ω) is evaluated by a mechanical rheometer in a destructive process by placing the specimen between two parallel plates and applying a sinusoidal shear strain, ε(ω), to the specimen. Here, ω is the angular oscillation frequency with the units of radians per second (rad/s). The consequent oscillatory stress induced within the sample, σ(ω), has both an in-phase component and an out-of-phase component with respect to ε(ω). The rheometer senses the torque and the displacement of the plates and retrieves the stress and strain magnitudes. Subsequently, G*(ω) is measured by calculating the ratio of applied stress to the resulting strain, i.e. σ(ω)/ε(ω), over a small frequency range. The ratio of the in-phase component of stress to the strain accounts for elastic (or storage) modulus. Besides, the ratio of stress component with 90-degree phase-lag with respect to strain represents the viscous (or loss) modulus. Alternatively, in the stress-controlled rheometers, a constant shear stress is applied to the specimen at various oscillation frequencies and the strain is evaluated. As for the case of strain-controlled rheometers, G* is evaluated by calculating the stress-strain ratio. Measuring the average bulk response over the entire sample volume in the rheometer precludes the inquiry of local mechanical heterogeneities and yields only the volume-averaged mechanical response of the specimen. Moreover, this type of mechanical test requires relatively large sample volumes and is not conducive for rare and precious biomaterials and tissue specimens.
The micro-scale mechanical properties of materials are conventionally probed using micro- and nano-indentation methods such as Atomic Force Microscopy (AFM)-based indentation, also known as force-mapping mode or force spectroscopy. The AFM force spectroscopy is a surface probing technique, capable of providing local elasticity maps of samples such as the ECM, live cells, and sub-cellular organelles. In force spectroscopy, a small micron-sized metallic tip is fixed in the proximity of a flexible cantilever. The cantilever is mounted on a high-precision piezoelectric stage, which controls the displacement of the tip with nano-meter precision. The pyramidal end of the tip may directly indent the sample. Alternatively, and more prevalently, a polystyrene bead of a few microns diameter may be glued at the tip end to increase the contact area.
Although the nominal value of the cantilever spring-constant is often provided by the manufacturer, it is imperative that a certain set of calibration steps be followed to re-evaluate the exact spring constant of the cantilever prior to measurements. This can be done, for instance, by thermal noise method via balancing of the ambient thermal energy of the cantilever environment with its free vibration and finding the resonance frequency. A Lorenzian model is fit to the resonance peak and the area under the curve is calculated as a measure of the resonance energy or tip deflection. By dividing the ambient thermal energy to the area under the curve, the exact spring constant of the cantilever is calculated.
Also vital to AFM operation is a calibration step that enables retrieving the cantilever deflection from the actual measured values, that is, detected voltage. In this procedure, the tip is lowered until it comes into physical contact with a hard calibration specimen, such as a glass slide. Since the hard calibration sample does not move, the deflection of the cantilever is equivalent to its displacement. This enables calculation of the conversion factor that relates the recorded voltage, the actual measured value, to the cantilever deflection.
After these cumbersome calibration steps, one may proceed to the force spectroscopy measurements of the specimen. As mentioned earlier, the cantilever is mounted on a precision z-stage. The z-stage is displaced so that the tip approaches, comes into physical contact, and presses further into the sample. As the tip pokes the specimen, the cantilever is deflected. To measure this deflection, a laser beam is focused at the extremities of the metallic tip. The reflection of the laser beam is monitored by a segmented photo-detector. Deflection of the cantilever stirs the laser beam, enabling the position-sensitive photo-detector to retrieve the deflection of the cantilever. Subsequently, Hooke's law is used to calculate the force as the product of cantilever deflection and spring constant. The force placed on the tip increases as it pokes and presses further into the sample, reaching a preset control value. Subsequently, the force is removed and the tip retracts. The z-stage displacement is plotted against the applied force during both the approach and retract phases. An appropriate model is fit to this force-distance curve to calculate the indentation modulus (E) of the specimen. This process needs to be repeated for each and every point on the sample within the region of interest. Towards this end, a pair of piezoelectric precision x-y stages translates either the cantilever or the sample to enable probing the entire region of interest. At each measurement point, the tip approaches towards and retracts back from the sample. Given that the approach-retract step takes at least a few seconds, the force-mapping AFM measurements are tediously time-consuming. Moreover, due to their contact-based nature, the indentation-based technologies are inherently invasive. In addition, the depth probed by the tip does not exceed a few microns, limiting AFM to a surface probe modality. Finally, the indentation modulus is merely reflecting the elastic behavior at a fixed indentation rate and does not probe the frequency-dependence or the dynamic viscosity.